Photonic crystal biosensor structure and fabrication method

ABSTRACT

The invention provides sensor compositions and method of making sensors.

GOVERNMENT INTERESTS

This invention was made with Government support under Grant NumberBES04-27657 awarded by the National Science Foundation. The Governmenthas certain rights in the invention.

BACKGROUND OF THE INVENTION:

Label-free optical sensors based upon surface structured photoniccrystals have recently been demonstrated as a highly sensitive methodfor performing a wide variety of biochemical and cell-based assays. See,e.g., Cunningham, et al., Label-Free Assays on the BIND System. Journalof Biomolecular Screening, 2004. 9:481-490. These sensors reflect only anarrow band of wavelengths when illuminated with white light at normalincidence, where positive shifts of the reflected peak wavelength value(PWV) indicate the adsorption of detected material on the sensorsurface. See, e.g., Cunningham, et al., Colorimetric resonant reflectionas a direct biochemical assay technique. Sensors and Actuators B, 2002.81:316-328. By spatially confining incident photons at the resonantwavelength, a high optical field is generated at the sensor surface thatextends a short distance into a test sample, much like an evanescentfield. The high degree of spatial confinement of resonant photons withinthe device structure leads to a strong interaction between the structureand adsorbed biomaterial, and to the ability to perform high resolutionimaging of protein and cell attachment. See, e.g., Li, et al., A newmethod for label-free imaging of biomolecular interactions. Sensors andActuators B, 2004. 99:6-13.

Previously, photonic crystal optical biosensors have been fabricatedfrom continuous sheets of plastic film using a process in which theperiodic surface structure is replicated from a silicon master waferusing a UV-cured polymer material. See, e.g., Cunningham, et al., Aplastic calorimetric resonant optical biosensor for multiparalleldetection of label-free biochemical interactions. Sensors and ActuatorsB, 2002. 85:219-226. This patterned polymer can be subsequently coatedwith a high refractive index TiO₂ layer that is generally thinner thanthe height of the surface structure. Such devices have been demonstratedfor a wide variety of biochemical and cell-based assays, with a massdensity sensitivity resolution less than 0.1 pg/mm² and a large dynamicrange enabling single cell detection. See, e.g., Lin et al., Alabel-free biosensor-based cell attachment assay for characterization ofcell surface molecules. Sensors and Actuators B, Accepted April 2005. Ingeneral, optimization of device sensitivity requires increasing theinteraction of the electromagnetic field intensity distribution with themolecules deposited atop the photonic crystal surface. Therefore,selection of optical materials and design of the surface structuretopology is aimed at extending the electromagnetic field profile fromthe interior regions of the photonic crystal (where they cannot interactwith adsorbed material) to the region adjacent to the photonic crystalthat includes the liquid test sample.

Methods are needed in the art to increase the sensitivity of these andother types of sensors and to decrease the cost of their manufacture.

SUMMARY OF THE INVENTION

One embodiment of the invention provides a method of making a sensor.The method comprises putting a liquid elastomer into a sub-wavelengthperiod grating structure mold and curing the liquid elastomer to a forma negative mold; removing the negative mold from the mold; setting thenegative mold into an uncured nanoporous film, wherein the uncurednanoporous film is supported by a substrate; allowing the nanoporousfilm to fully cure, partially cure or not cure; removing the negativemold from the fully cured, partially cured, or uncured nanoporous film;depositing a high dielectric constant dielectric material onto the fullycured, partially cured, or uncured nanoporous film, wherein a sensor isformed. The nanoporous film can comprise porous silica xerogel, porousaerogels, porous hydrogen silsesquioxane, a B staged polymer, porousmethyl silsesquioxane, porous poly(arylene ether), or combinationsthereof. The substrate can comprise plastic, epoxy, or glass. The highdielectric constant dielectric material can be tin oxide, tantalumpentoxide, zinc sulfide, titanium dioxide, silicon nitride, or acombination thereof. The refractive index of the high dielectricconstant dielectric material can be from about 1.8 to about 3.0. Themethod can further comprise depositing a cover layer on the highdielectric constant dielectric material. The method can further compriseimmobilizing one or more specific binding substances on the surface ofthe high dielectric constant dielectric material. The method can furthercomprise immobilizing one or more specific binding substances on thesurface of the cover layer. The one or more specific binding substancescan be detection label-free. The refractive index of the porous film canbe from about 1.1 to about 2.2. The refractive index of the porous filmcan be from about 1.1 to about 1.5. The period of the sub-wavelengthperiod grating structure can be from about 200 nm to about 1,500 nm andthe depth of the sub-wavelength period grating structure can be fromabout 50 nm to about 900 nm. The refractive index of the substrate canbe about 1.4 to about 1.6. The thickness of the high dielectric constantdielectric material can be about 30 nm to about 700 nm and the thicknessof the nanoporous material can be about 10 mn to about 5,000 nm.

Another embodiment of the invention provides a method of making asensor. The method comprises curing a layer of nanoporous materialhaving a low refractive index onto a substrate; depositing a highdielectric constant dielectric material on top of the nanoporousmaterial; and patterning the high dielectric constant dielectricmaterial into a grating. The nanoporous film can comprise porous silicaxerogel, porous aerogels, porous hydrogen silsesquioxane, a B stagedpolymer, porous methyl silsesquioxane, porous poly(arylene ether), orcombinations thereof. The substrate can comprise plastic, epoxy, orglass. The high dielectric constant dielectric material can be tinoxide, tantalum pentoxide, zinc sulfide, titanium dioxide, siliconnitride, or a combination thereof. The refractive index of the highdielectric constant dielectric material is about 1.8 to about 3.0. Themethod can further comprise depositing a cover layer on the highrefractive index dielectric material. The method can further compriseimmobilizing one or more specific binding substances on the surface ofthe high refractive index dielectric material. The method can furthercomprise immobilizing one or more specific binding substances on thesurface of the cover layer. The one or more specific binding substancescan be detection label-free. The refractive index of the nanoporous filmcan be from about 1.1 to about 2.2. The refractive index of thenanoporous film can be from about 1.1 to about 1.5. The period of thesub-wavelength period grating structure can be about 200 nm to about1,500 nm and the depth of the sub-wavelength period grating structurecan be about 50 nm to about 900 nm. The refractive index of thesubstrate can be about 1.4 to about 1.6. The thickness of the highdielectric constant dielectric material can be about 30 nm to about 700nm and the thickness of the nanoporous material can be about 10 mn toabout 5,000 nm.

Therefore, use of an extremely low refractive index material for thesurface structure in sensors substantially increases detectionsensitivity. Therefore, substances can be measured in test samples with2-4× lower concentrations, molecular weights, or binding affinities thanhave been possible previously.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a schematic of a nano-replicated nanoporous photoniccrystal biosensor.

FIG. 2A-B shows (A) bulk and (B) surface shift predicted by GSolversimulations.

FIG. 3A-B shows SEM images of imprinted and cured periodic NANOGLASS®structure.

FIG. 4 shows the experimental response of nanoporous sensor underimmersion in de-ionized water (DI) and isopropyl alcohol (IPA).

FIG. 5A-F shows process flow for nanoporous sensor fabrication.

FIG. 6 shows a schematic of high dielectric constant nanoporous photoniccrystal sensor.

FIG. 7 shows a cross-section schematic of porous glass sensor.

FIG. 8 shows resonant peak of porous glass sensor exposed to deionizedwater, as predicted by RCWA simulation.

FIG. 9 shows experimentally measured resonant peak of nanoporous glasssensor immersed in deionized water.

FIG. 10 shows a kinetic plot comparing PWV shifts for PPL deposited ontoboth porous glass and polymer sensor designs.

FIG. 11 shows a partial profile of PWV shift versus polymer thickness,where alternating layers of PSS and PAH contribute to the total measuredshift.

FIG. 12 shows a kinetic plot comparing PWV shifts for protein Adeposited onto both porous glass and polymer sensor designs.

FIG. 13 shows binding kinetics of three animal IgGs to protein Ameasured with a nanoporous glass sensor.

FIG. 14 shows sensor comparison of PWV shifts for each of the differentIgG-protein A interactions.

DETAILED DESCRIPTION OF THE INVENTION

One embodiment of the invention provides a sensor that can be used to,inter alia, detect organic or inorganic material, such as protein, DNA,small molecules, viruses, cells, and bacteria, without the requirementof a label, such as fluorescent or radioactive labels. Photonic crystalsensors of the invention reflect only a very narrow band of wavelengthsor one wavelength when illuminated with a broad wavelength light source(such as a white light or LED). The reflected color shifts to longerwavelengths in response to attachment of material to the sensor surface.Photonic crystal sensor structures of the invention provide 2-4× highersensitivity than previously described structures. A key difference inthe sensor structure that provides higher sensitivity is the replacementof a polymer sub-wavelength period grating structure with a nanoporouslow refractive index material.

Methods for fabricating sensor structures that enable low costmanufacturing are also disclosed. Sensor structures of the inventionhave higher sensitivity than previous structures due to the use ofnanoporous low refractive index material instead of a polymersub-wavelength period grating structure. When the refractive index ofthe sensor structure directly beneath (and alternatively including) thesub-wavelength grating structure is reduced below the refractive indexof any liquid used in a sample, the electromagnetic field of thephotonic crystal interacts more strongly with the test sample, yieldinga structure whose reflected wavelength is more strongly tuned by a givenamount of adsorbed biological material. The system is capable ofdetecting, e.g., a single cell attached to its surface.

The principles of the instant invention can also be applied to, e.g.,evanescent wave-based biosensors and any biosensors incorporating anoptical waveguide. See, e.g., U.S. Pat. Nos. 4,815,843; 5,071,248;5,738,825.

The sensors have utility in, inter alia, the fields of pharmaceuticalresearch (e.g., high throughput screening, secondary screening, qualitycontrol, cytotoxicity, clinical trial evaluation), life science research(e.g., proteomics, protein interaction analysis, DNA-protein interactionanalysis, enzyme-substrate interaction analysis, cell-proteininteraction analysis), diagnostic tests (e.g., protein presence, cellidentification), and environmental detection (bacterial and sporedetection and identification). Previous patent applications andpublications describe how the photonic crystal biosensor surface, incombination with a high resolution imaging instrument, can be used as aplatform for performing many biochemical assays in parallel upon onsingle surface, using only nanoliters of sample material. See, e.g.,U.S. Pat. Publ. Nos.: 2002/0168295; 2002/0127565; 2004/0132172;2004/0151626; 2003/0027328; 2003/0027327; 2003/017581; 2003/0068657;2003/0059855; 2003/0113766; 2003/0092075; 2003/0026891; 2003/0026891;2003/0032039; 2003/0017580; 2003/0077660; 2004/0132214.

Photonic Crystal Sensors

A photonic crystal sensor of the invention can be used to create a sharpoptical resonant reflection at a particular wavelength that can be usedto track with high sensitivity the interaction of molecules, such asbiological materials.

Photonic crystal sensors comprise a subwavelength structured surface.Subwavelength structured surfaces are a type of diffractive optic thatcan mimic the effect of thin-film coatings. See, e.g., Peng & Morris,“Resonant scattering from two-dimensional gratings,” J. Opt. Soc. Am. A,Vol. 13, No. 5, p. 993, May 1996; Magnusson, & Wang, “New principle foroptical filters,” Appl. Phys. Lett., 61, No. 9, p. 1022, August, 1992;Peng & Morris, “Experimental demonstration of resonant anomalies indiffraction from two-dimensional gratings,” Optics Letters, Vol. 21, No.8, p. 549, April, 1996. A grating of a photonic crystal sensor of theinvention has a grating period that is small compared to the wavelengthof incident light such that no diffractive orders other than thereflected and transmitted zeroth orders are allowed. A photonic crystalsensor can comprise a grating, which is comprised of or coated with ahigh dielectric constant dielectric material, sandwiched between asubstrate layer and a cover layer that fills the grating grooves.Optionally, a cover layer is not used. The grating structure selectivelycouples light at a narrow band of wavelengths. This highly sensitivecoupling condition can produce a resonant grating effect on thereflected radiation spectrum, resulting in a narrow band of reflected ortransmitted wavelengths. The depth and period of the grating are lessthan the wavelength of the resonant grating effect.

The reflected or transmitted color of a photonic crystal sensorstructure can be modified by the addition of molecules such as specificbinding substances or binding partners or both to the upper surface ofthe cover layer or the grating surface. The added molecules increase theoptical path length of incident radiation through the sensor structure,and thus modify the wavelength at which maximum reflectance ortransmittance will occur.

In one embodiment, a sensor, when illuminated with white light, isdesigned to reflect only a single wavelength or a narrow band ofwavelengths. When molecules are attached to the surface of the sensor,the reflected wavelength (color) is shifted due to the change of theoptical path of light that is coupled into the grating. By immobilizingmolecules, such as specific binding substances to a sensor surface,complementary binding partner molecules can be detected without the useof any kind of fluorescent probe or particle label. The detectiontechnique can be performed with the sensor surface either immersed influid or dried.

When a photonic crystal sensor is illuminated with collimated whitelight and reflects only a narrow band of wavelengths, or a single bandof wavelengths is reflected. The narrow wavelength band is described asa wavelength “peak.” The “peak wavelength value” (PWV) changes whenmolecules are deposited or removed from the sensor surface. A readoutinstrument illuminates distinct locations on the sensor surface withcollimated white light, and collects collimated reflected light. Thecollected light is gathered into a wavelength spectrometer fordetermination of PWV.

FIG. 1 shows a structure of a photonic crystal sensor of the invention.The sensor comprises a substrate, a patterned, low-k, nanoporousmaterial, and a substantially uniform, high refractive index coating.The surface of the low-k nanoporous material is patterned into asub-wavelength period grating structure onto which the high refractiveindex material is deposited.

In general, a low-k dielectric material of the invention has adielectric constant, k, of about 1.1 to about 3.9. Examples of low-kdielectric materials include, for example: fluorosilicate glass (about3.2-about 3.9); polyimides (about 3.1-about 3); hydrogen silsesquioxane(HSQ) (about 2.9-about 3.2); diamond-like carbon (about 2.7-about 3.4);black diamond (SiCOH) (about 2.7-about 3.3); parylene-N (about 2.7);B-staged polymers (CYCLOTENE™ and SiLK™) (about 2.6-about 2.7);fluorinated polyimides (about 2.5-about 2.9); methyl silsequioxane (MSQ)(about 2.6-about 2.8); poly(arylene ether) (PAE) (about 2.6-about 2.8);fluorinated DLC (about 2.4-about 2.8); parylene-F (about 2.4-about 2.5);PTFE (about 1.9); porous silica xerogels and aerogels (about 1.1-about2.2); porous hydrogen silsesquioxane (HSQ) (about 1.7-about 2.2); porousSiLK™ (a B staged polymer) (about 1.5-about 2.0); porous methylsilsesquioxane (MSQ) (about 1.8-about 2.2); porous poly(arylene ether)(PAE) (about 1.8-about 2.2).

A low-k nanoporous material is an inorganic, porous, oxide-like lowdielectric material, wherein the refractive index, n, is about 1.1 toabout 2.2, and preferably about 1.1 to about 1.5. A low-k nanoporousmaterial can be, for example, porous silica xerogels and aerogels (about1.1-about 2.2); porous HSQ (about 1.7-about 2.2); porous SiLK™ (a Bstaged polymer) (about 1.5-about 2.0); porous MSQ (about 1.8-about 2.2);porous PAE (about 1.8-about 2.2). In one embodiment of the invention thenanoporous material is NANOGLASS®, which is porous SiO₂. Porosity iscreated in the SiO₂ thereby reducing the dielectric constant from about3.9 to as low as 1.9.

A material with a high refractive index, suitable for the inventionincludes, e.g., tin oxide, tantalum pentoxide, zinc sulfide, titaniumdioxide, silicon nitride, or a combination thereof. A high k dielectricmaterial has a refractive index of about 1.8 to about 3.0. Refractiveindex, n, describes the optical characteristics of a medium and isdefined as the ratio of the speed of light in free space over the speedof light in the medium. A substrate can comprise, for example, glass,plastic or epoxy.

In one embodiment of the invention a sensor is defined by the followingparameters:

n_(hiK) About 1.8 to about 3.0 n_(nano) About 1.1 to about 1.5 n_(sub)About 1.4 to about 1.6 Λ About 200 nm to about 1500 nm d About 50 nm toabout 900 nm t_(hiK) About 30 nm to about 700 nm t_(nano) About 10 nm toabout 5000 nm

In another embodiment of the invention, the sensor structure comprisesthe following materials:

Substrate Material Glass Nanoporous Material Nanoglass ® (HoneywellInternational, Santa Clara, CA) High Refractive Index TiO₂ Coatingand is defined by the following parameters.

n_(hiK) 2.25 n_(nano) 1.17 n_(sub) 1.50 Λ 550 nm d 170 nm t_(hiK) 120 nmt_(nano) 600 nm

Simulations of the above embodiment were performed using GSolver(Grating Solver Development Co., Allen, Tex.) and FDTD Solutions(Lumerical Solutions, Inc., Vancouver, BC, Canada). The results shown inFIG. 2A predict the bulk sensitivity to be improved by more than afactor of two over that of previous designs. Bulk sensitivity isdetermined by the bulk shift coefficient, defined and calculated forthis embodiment below.

$\begin{matrix}{\frac{\Delta\;{PWV}}{\Delta\; n} = {\frac{\lambda_{IPA} - \lambda_{DI}}{n_{IPA} - n_{DI}} = {\frac{794.4 - 779.6}{1.378 - 1.330} = 308.3}}} & (1)\end{matrix}$Both simulation and experimental data for designs that do notincorporate a nanoporous material give bulk shift coefficients ofapproximately 150.

Since the proposed device functions by evanescent field interactionswith materials very near the sensor surface, it is instructive toconsider a refractive index shift not only of the entire bulk media butalso of a thin layer atop the sensor. FIG. 2B shows GSOLVER simulationresults with a 20 nm thick “biological coating” modeled by a layer witha refractive index of 1.40. While individual biological molecules orfractions of biological molecule monolayers do not have a definedrefractive index value, the biological layer was modeled as a uniformthin film of defined thickness for the sake of illustration.

The SEM images of the patterned NANOGLASS® structure shown in FIG. 3give evidence of a successful imprinting process. Upon deposition ofTiO₂, the sensitivity of the completed sensor was interrogated usingde-ionized water and isopropyl alcohol by examining the resulting peakwavelength (PWV) shift captured with a spectrometer on the readoutinstrumentation. Applying Equation 1 using the experimental data fromFIG. 4, the bulk shift coefficient can be calculated as:

$\frac{\Delta\;{PWV}}{\Delta\; n} = {\frac{\lambda_{IPA} - \lambda_{DI}}{n_{IPA} - n_{DI}} = {\frac{855.1 - 841.1}{1.378 - 1.330} = 291.7}}$which agrees to within 5% of that demonstrated though simulation.

A cross-sectional profile of a subwavelength grating can comprise anyperiodically repeating function, for example, a “square-wave.” A gratingcan be comprised of a repeating pattern of shapes such as continuousparallel lines, squares, circles, ellipses, triangles, trapezoids,sinusoidal waves, ovals, rectangles, and hexagons.

A sensor can comprise a one-dimensional linear grating surfacestructure, i.e., a series of parallel lines or grooves. While atwo-dimensional grating has features in two lateral directions acrossthe plane of the sensor surface that are both subwavelength, thecross-section of a one-dimensional grating is only subwavelength in onelateral direction, while the long dimension can be greater thanwavelength of the resonant grating effect. These include, for example,triangular or v-shaped, u-shaped, upside-down v- or u-shapes,sinusoidal, trapezoidal, stepped and square. The grating can also besinusoidally varying in height.

An alternate sensor structure can be used that consists of a set ofconcentric rings. In this structure, the difference between the insidediameter and the outside diameter of each concentric ring is equal toabout one-half of a grating period. Each successive ring has an insidediameter that is about one grating period greater than the insidediameter of the previous ring. The concentric ring pattern extends tocover a single sensor location—such as a microarray spot or a microtiterplate well. Each separate microarray spot or microtiter plate well has aseparate concentric ring pattern centered within it. All polarizationdirections of such a structure have the same cross-sectional profile.The grating period of a concentric ring structure is less than thewavelength of the resonantly reflected light

A sensor of the invention can further comprise a cover layer on thesurface of a grating opposite to a substrate layer. Where a cover layeris present, the one or more specific binding substances are immobilizedon the surface of the cover layer opposite to the grating. Preferably, acover layer comprises a material that has a lower refractive index thana material that comprises the grating. A cover layer can be comprisedof, for example, glass (including spin-on glass (SOG)), epoxy, orplastic.

Resonant reflection can also be obtained without a planarizing coverlayer over the grating. Without the use of a planarizing cover layer,the surrounding medium (such as air or water) fills the grating.Therefore, molecules are immobilized to the sensor on all surfaces of agrating exposed to the molecules, rather than only on an upper surface.

The invention provides resonant reflection structures and transmissionfilter structures. For a resonant reflection structure, light output ismeasured on the same side of the structure as the illuminating lightbeam. For a transmission filter structure, light output is measured onthe opposite side of the structure as the illuminating beam. Thereflected and transmitted signals are complementary. That is, if awavelength is strongly reflected, it is weakly transmitted. Assuming noenergy is absorbed in the structure itself, the reflected+transmittedenergy at any given wavelength is constant. The resonant reflectionstructures and transmission filters are designed to give a highlyefficient reflection at a specified wavelength. Thus, a reflectionfilter will “pass” a narrow band of wavelengths, while a transmissionfilter will “cut” a narrow band of wavelengths from incident light.

In one embodiment of the invention, an optical device is provided. Anoptical device comprises a structure similar to any sensor of theinvention; however, an optical device does not comprise one or morebinding substances immobilized on the grating. An optical device can beused as a narrow band optical filter.

Evanescent wave-based sensors can comprise a waveguiding film supportedby a substrate; between the waveguiding film (and optionally as part ofthe substrate) is a diffraction grating. See, e.g., U.S. Pat. No.4,815,843. A low-k dielectric material, such as low-k nanoporousmaterial can be used for the diffraction grating or the combined low-knanoporous material and substrate. The waveguide comprises waveguidingfilm and the substrate. The waveguiding film can be, e.g., tin oxide,tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, ora combination thereof, or a polymer such as polystryrole orpolycarbonate. A diffraction grating exists at the interface of thewaveguiding film and the substrate or in the volume of the waveguidingfilm. The diffraction grating comprises a low-k material, such as low-knanoporous material. The refractive index of the waveguiding film ishigher than the index of the adjacent media (i.e., the substrate and thetest sample). The substrate can be, e.g., plastic, glass or epoxy. Aspecific binding substance can be immobilized on the surface of thewaveguiding film and a test sample added to the surface. Laser lightpropagates in the waveguiding film by total internal reflection. Changesin refractive index of the waveguiding film caused by molecules bindingto it can be detected by observing changes in the angle of the emitted,out-coupled light.

Production of Sensors

Sensors of the invention can be produced using a flexible rubbertemplate for embossing the grating structure into the nanoporousmaterial while the nanoporous material is in an uncured, deformablestate. Unlike nonflexible solid templates, the flexible rubber templateallows solvent vapors, generated by the nanoporous material's curingprocess, to escape. Many flexible templates can be generated from asingle silicon wafer “master” template at low cost, and a singleflexible template can be used multiple times to inexpensively producemany structured nanoporous sub-wavelength grating structures.

Sensors can be produced inexpensively over large surface areas and canalso be, for example, incorporated into single-use standard disposableassay liquid handling formats such as microplates, microarray slides, ormicrofluidic chips.

A process flow for fabricating a photonic crystal incorporating ananoporous layer is outlined in FIG. 5. A patterned “master” wafer,usually silicon or glass, which contains features that will correspondprecisely with those later imprinted into the porous film is designed(see FIG. 5A). The master is then used as a mold into which a liquidelastomer is poured, as shown in (FIG. 5B). Upon curing, the newlyformed negative rubber “daughter” mold is carefully peeled away from themaster. After application of the porous film to the desired substrate,the daughter mold is set atop the uncured film, e.g., as depicted in(FIG. 5D). With the mold in place, the porous material is partiallycured, fully cured or not cured. The gas-permeable rubber mold allowssolvent evaporation during this curing process. Once the film cansustain a rigid shape, the daughter mold is peeled away and theremaining structure is allowed to fully cure. A completed deviceillustrated in (FIG. 5F) is obtained by depositing a thin highrefractive index material uniformly across the patterned surface of theporous film.

Another approach for fabrication of a photonic crystal biosensorincorporating a nanoporous layer is illustrated in FIG. 6. With thisstructure, a layer of nanoporous material is cured onto a substrate.Next, a high dielectric constant material is uniformly deposited on topof the porous layer. A high dielectric constant material has adielectric constant, k, greater than about 5% higher than the k of thenanoporous material. In one embodiment of the invention the highdielectric constant material has a k of greater than about 3.5. Thedeposited high-k material is then patterned by e-beam or DUVlithography, and subsequently etched to obtain the desired features.While this sensor design is not as cost effective due to the need forhigh-resolution lithographic processes for each device, it shows promisefor obtaining sensitivity enhancements similar to those seen with theaforementioned sensor fabricated by imprinting.

Evanescent-wave based biosensors can also be made using the sameprocesses as described herein.

Specific Binding Substances and Binding Partners

One or more specific binding substances can be immobilized on a gratingor cover layer, if present, by for example, physical adsorption or bychemical binding. A specific binding substance can be, for example, anorganic molecule, such as a nucleic acid, polypeptide, antigen,polyclonal antibody, monoclonal antibody, single chain antibody (scFv),F(ab) fragment, F(ab′)₂ fragment, Fv fragment, small organic molecule,cell, virus, bacteria, polymer, peptide solutions, single- ordouble-stranded DNA solutions, RNA solutions, solutions containingcompounds from a combinatorial chemical library, or biological sample;or an inorganic molecule. A biological sample can be for example, blood,plasma, serum, gastrointestinal secretions, homogenates of tissues ortumors, synovial fluid, feces, saliva, sputum, cyst fluid, amioticfluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen,lymphatic fluid, tears, or prostatitc fluid.

Preferably, one or more specific binding substances are arranged in amicroarray of distinct locations on a sensor. One or more specificbinding substances can be bound to their specific binding partners. Amicroarray of specific binding substances comprises one or more specificbinding substances on a surface of a sensor of the invention such that asurface contains many distinct locations, each with a different specificbinding substance or with a different amount of a specific bindingsubstance. For example, an array can comprise 1, 10, 100, 1,000, 10,000,or 100,000 distinct locations. Such a sensor surface is called amicroarray because one or more specific binding substances are typicallylaid out in a regular grid pattern in x-y coordinates. However, amicroarray of the invention can comprise one or more specific bindingsubstances laid out in any type of regular or irregular pattern. Forexample, distinct locations can define a microarray of spots of one ormore specific binding substances. A microarray spot can be about 50 toabout 500 microns in diameter. A microarray spot can also be about 150to about 200 microns in diameter.

A microarray on a sensor of the invention can be created by placingmicrodroplets of one or more specific binding substances onto, forexample, an x-y grid of locations on a grating or cover layer surface.When the sensor is exposed to a test sample comprising one or morebinding partners, the binding partners will be preferentially attractedto distinct locations on the microarray that comprise specific bindingsubstances that have high affinity for the binding partners. Some of thedistinct locations will gather binding partners onto their surface,while other locations will not.

A specific binding substance specifically binds to a binding partnerthat is contacted with the surface of a sensor of the invention. Aspecific binding substance specifically binds to its binding partner,but does not substantially bind other binding-partners contacted withthe surface of a sensor. For example, where the specific bindingsubstance is an antibody and its binding partner is a particularantigen, the antibody specifically binds to the particular antigen, butdoes not substantially bind other antigens. A binding partner can be,for example, a nucleic acid, polypeptide, antigen, polyclonal antibody,monoclonal antibody, single chain antibody (scFv), F(ab) fragment,F(ab′)₂ fragment, Fv fragment, small organic molecule, cell, virus,bacteria, polymer, peptide solutions, single- or double-stranded DNAsolutions, RNA solutions, solutions containing compounds from acombinatorial chemical library, an inorganic molecule, or a biologicalsample.

One example of a microarray of the invention is a nucleic acidmicroarray, in which each distinct location within the array contains adifferent nucleic acid molecule. In this embodiment, the spots withinthe nucleic acid microarray detect complementary chemical binding withan opposing strand of a nucleic acid in a test sample.

While microtiter plates are the most common format used for biochemicalassays, microarrays are increasingly seen as a means for maximizing thenumber of biochemical interactions that can be measured at one timewhile minimizing the volume of precious reagents. By application ofspecific binding substances with a microarray spotter onto a sensor ofthe invention, specific binding substance densities of 10,000 specificbinding substances/in² can be obtained. By focusing an illumination beamto interrogate a single microarray location, a sensor can be used as alabel-free microarray readout system.

While it is not necessary for specific binding substances or bindingpartners to comprise a detectable label, detectable labels can be usedto detect specific binding substances or binding partners on the surfaceof a sensor. Where specific binding substances and binding partners ofthe instant invention are free of detection labels, they can stillcomprise other types of labels and markers for enhancement of assaysensitivity, immobilization of specific binding partners to a biosensorsurface, enhancement of binding or hybridization of specific bindingsubstances to their binding partners, and for other purposes.

Immobilization of One or More Specific Binding Substances

Molecules can be immobilized onto a sensor is so that they will not bewashed away by rinsing procedures, and so that binding to molecules in atest sample is unimpeded by the sensor surface. Several different typesof surface chemistry strategies have been implemented for covalentattachment of molecules to, for example, glass for use in various typesof microarrays and sensors. These same methods can be readily adapted toa sensor of the invention.

One or more types of molecules can be attached to a sensor surface byphysical adsorption (i.e., without the use of chemical linkers) or bychemical binding (i.e., with the use of chemical linkers). Chemicalbinding can generate stronger attachment of molecules on a sensorsurface and provide defined orientation and conformation of thesurface-bound molecules.

Other types of chemical binding include, for example, amine activation,aldehyde activation, and nickel activation. These surfaces can be usedto attach several different types of chemical linkers to a sensorsurface. While an amine surface can be used to attach several types oflinker molecules, an aldehyde surface can be used to bind proteinsdirectly, without an additional linker. A nickel surface can be used tobind molecules that have an incorporated histidine (“his”) tag.Detection of “his-tagged” molecules with a nickel-activated surface iswell known in the art (Whitesides, Anal. Chem. 68, 490, (1996)).

Immobilization of specific binding substances to plastic, epoxy, or highrefractive index material can be performed essentially as described forimmobilization to glass. However, an acid wash step can be eliminatedwhere such a treatment would damage the material to which the specificbinding substances are immobilized.

Liquid-Containing Vessels

A sensor of the invention can comprise an inner surface, for example, abottom surface of a liquid-containing vessel. A liquid-containing vesselcan be, for example, a microtiter plate well, a test tube, a petri dish,or a microfluidic channel. One embodiment of this invention is a sensorthat is incorporated into any type of microtiter plate. For example, asensor can be incorporated into the bottom surface of a microtiter plateby assembling the walls of the reaction vessels over the resonantreflection surface, so that each reaction “spot” can be exposed to adistinct test sample. Therefore, each individual microtiter plate wellcan act as a separate reaction vessel. Separate chemical reactions can,therefore, occur within adjacent wells without intermixing reactionfluids, and chemically-distinct test solutions can be applied-toindividual wells.

The most common assay formats for pharmaceutical high-throughputscreening laboratories, molecular biology research laboratories, anddiagnostic assay laboratories are microtiter plates. The plates arestandard-sized plastic cartridges that can contain 96, 384, or 1536individual reaction vessels arranged in a grid. Due to the standardmechanical configuration of these plates, liquid dispensing, roboticplate handling, and detection systems are designed to work with thiscommon format. A sensor of the invention can be incorporated into thebottom surface of a standard microtiter plate. Because the sensorsurface can be fabricated in large areas, and because the readout systemdoes not make physical contact with the sensor surface, an arbitrarynumber of individual sensor areas can be defined that are only limitedby the focus resolution of the illumination optics and the x-y stagethat scans the illumination/detection probe across the sensor surface.

A sensor can also be incorporated into other disposable laboratory assayformats, such as microarray slides, flow cells, and cell culture plates.Incorporation of a sensor into common laboratory formats is desirablefor compatibility with existing microarray handling equipment such asspotters and incubation chambers.

Methods of Using Sensors

Sensors of the invention can be used, e.g., study one or a number ofmolecule/molecule interactions in parallel; for example, binding of oneor more specific binding substances to their respective binding partnerscan be detected, without the use of labels, by applying one or morebinding partners to a sensor that has one or more specific bindingsubstances immobilized on its surface. A sensor is illuminated withlight and a maximum in reflected wavelength, or a minimum in transmittedwavelength of light is detected from the sensor. If one or more specificbinding substances have bound to their respective binding partners, thenthe reflected wavelength of light is shifted as compared to a situationwhere one or more specific binding substances have not bound to theirrespective binding partners. Where a sensor is coated with an array ofdistinct locations containing the one or more specific bindingsubstances, then a maximum in reflected wavelength or minimum intransmitted wavelength of light is detected from each distinct locationof the sensor.

In one embodiment of the invention, a variety of specific bindingsubstances, for example, antibodies, can be immobilized in an arrayformat onto a sensor of the invention. The sensor is then contacted witha test sample of interest comprising binding partners, such as proteins.Only the proteins that specifically bind to the antibodies immobilizedon the sensor remain bound to the sensor. Such an approach isessentially a large-scale version of an enzyme-linked immunosorbentassay; however, the use of an enzyme or fluorescent label is notrequired.

The activity of an enzyme can be detected by detecting the reflectedwavelength of light from a sensor on which one or more specific bindingsubstances have been immobilized and applying one or more enzymes to thesensor. The sensor is washed and illuminated with light. The reflectedwavelength of light is detected from the sensor. Where the one or moreenzymes have altered the one or more specific binding substances of thesensor by enzymatic activity, the reflected wavelength of light isshifted.

Additionally, a test sample, for example, cell lysates containingbinding partners can be applied to a sensor of the invention, followedby washing to remove unbound material. The binding partners that bind toa sensor can subsequently be eluted from the sensor and identified by,for example, mass spectrometry. Optionally, a phage DNA display librarycan be applied to a sensor of the invention followed by washing toremove unbound material. Individual phage particles bound to the sensorcan be isolated and the inserts in these phage particles can then besequenced to determine the identities of the binding partners.

For the above applications, and in particular proteomics applications,the ability to selectively bind material, such as binding partners froma test sample onto a sensor of the invention, followed by the ability toselectively remove bound material from a distinct location of the sensorfor further analysis is advantageous. Sensors of the invention are alsocapable of detecting and quantifying the amount of a binding partnerfrom a sample that is bound to a sensor array distinct location bymeasuring the shift in reflected wavelength of light. Additionally, thewavelength shift at one distinct sensor location can be compared topositive and negative controls at other distinct sensor locations todetermine the amount of a binding partner that is bound to a sensorarray distinct location.

In one embodiment of the invention, an interaction of a first moleculewith a second test molecule can be detected. A sensor as described aboveis used; however, there are no specific binding substances immobilizedon its surface. Therefore, the sensor comprises a one- ortwo-dimensional grating, a substrate layer that supports the one- ortwo-dimensional grating, and optionally, a cover layer. As describedabove, when the sensor is illuminated a resonant grating effect isproduced on the reflected radiation spectrum, and the depth and periodof the grating are less than the wavelength of the resonant gratingeffect.

To detect an interaction of a first molecule with a second testmolecule, a mixture of the first and second molecules is applied to adistinct location on a sensor. A distinct location can be one spot orwell on a sensor or can be a large area on a sensor. A mixture of thefirst molecule with a third control molecule is also applied to adistinct location on a sensor. The sensor can be the same sensor asdescribed above, or can be a second sensor. If the sensor is the samesensor, a second distinct location can be used for the mixture of thefirst molecule and the third control molecule. Alternatively, the samedistinct sensor location can be used after the first and secondmolecules are washed from the sensor. The third control molecule doesnot interact with the first molecule and is about the same size as thefirst molecule. A shift in the reflected wavelength of light from thedistinct locations of the sensor or sensors is measured. If the shift inthe reflected wavelength of light from the distinct location having thefirst molecule and the second test molecule is greater than the shift inthe reflected wavelength from the distinct location having the firstmolecule and the third control molecule, then the first molecule and thesecond test molecule interact. Interaction can be, for example,hybridization of nucleic acid molecules, specific binding of an antibodyor antibody fragment to an antigen, and binding of polypeptides. A firstmolecule, second test molecule, or third control molecule can be, forexample, a nucleic acid, polypeptide, antigen, polyclonal antibody,monoclonal antibody, single chain antibody (scFv), F(ab) fragment,F(ab′)₂ fragment, Fv fragment, small organic molecule, cell, virus, andbacteria.

All patents, patent applications, and other scientific or technicalwritings referred to anywhere herein are incorporated by reference intheir entirety. The methods and compositions described herein aspresently representative of preferred embodiments are exemplary and arenot intended as limitations on the scope of the invention. Changestherein and other uses will be evident to those skilled in the art, andare encompassed within the spirit of the invention. The inventionillustratively described herein suitably can be practiced in the absenceof any element or elements, limitation or limitations that are notspecifically disclosed herein. Thus, for example, in each instanceherein any of the terms “comprising”, “consisting essentially of”, and“consisting of” can be replaced with either of the other two terms. Theterms and expressions which have been employed are used as terms ofdescription and not of limitation, and there is no intention in the useof such terms and expressions of excluding any equivalents of thefeatures shown and described or portions thereof, but it is recognizedthat various modifications are possible within the scope of theinvention claimed. Thus, it should be understood that although thepresent invention has been specifically disclosed by embodiments andoptional features, modification and variation of the concepts hereindisclosed are considered to be within the scope of this invention asdefined by the description and the appended claims.

In addition, where features or aspects of the invention are described interms of Markush groups or other grouping of alternatives, those skilledin the art will recognize that the invention is also thereby describedin terms of any individual member or subgroup of members of the Markushgroup or other group.

EXAMPLES Example 1 Computer Simulation

Rigorous Coupled Wave Analysis (RCWA) and Finite Difference Time Domain(FDTD) simulations were used to predict the resonant wavelength and bulkrefractive index sensitivity of a one-dimensional surface photoniccrystal biosensor. The device incorporates a low-index (n=1.17)nanoporous dielectric surface structure in place of the polymer (n=1.39)surface structure reported previously. A soft contact embossing methodwas used to create a surface-structured low-index porous film on glasssubstrates with a depth and period that are identical to the previouspolymer structures to enable a side-by-side sensitivity comparison. Thesensitivity of porous glass biosensors was compared to nonporous polymerbiosensors through methods that characterize sensitivity to bulkrefractive index and surface-adsorbed material. Finally, a proteinbinding assay comparison was performed to demonstrate sensor stabilityand the ability to functionalize the device for selective detection.

The polymer and porous glass sensors were modeled and simulated usingtwo software packages. First, a 2-D diffraction grating analysis tool(GSOLVER) employing the RCWA algorithm provides a quick and simplemethod for initial sensor modeling. Second, FDTD (Lumerical) provides amuch more versatile and powerful tool that can calculate any fieldcomponent at any temporal or spectral location for an arbitrary opticaldevice illuminated by an arbitrary source. See, e.g., Kunz & Luebbers,The Finite Difference Time Domain Method for Electromagnetics. 1993,Boca Raton: CRC Press. FDTD was used to verify RCWA results and to gaindeeper insight into the effects of modifying the sensor structure.

RCWA and FDTD simulations both indicated that replacement of thepatterned UV-cured polymer of previous devices with a material of lowerrefractive index would produce a two-fold increase in the bulk shiftcoefficient. The resonant wavelength of the porous glass sensor immersedin DI H₂O was predicted by RCWA to be 844.3 nm with a full-width athalf-maximum (FWHM) of approximately 2 nm, as shown in FIG. 8.Simulation predicts further improvements in the bulk shift coefficientwith slight modifications to the sensor geometry.

The bulk sensitivity test using DI H₂O and IPA was performed on 23porous glass sensors and 11 polymer sensors. The average PWV shifts were13.6±2.4 nm and 5.1±1.5 nm for the porous glass and polymer sensors,respectively. The bulk shift coefficient (ΔPWV/AΔ) of the porous glasssensor is measured to be 2.7±1.2 times greater than that of the polymerdevice. Measurements of the porous glass device in DI H₂O give anaverage PWV of 829.5±16.5 nm and FWHM of 3.5±2.5 nm. One of the measuredspectra is illustrated in FIG. 9, where the response has been normalizedto a perfect reflector to account for any instrumentation losses. Thelower reflection efficiency and broader FWHM measured from thereplicated devices can be attributed to small but measurable materiallosses and to imperfections observed in the replicated structure. Thelarge variability of measured spectral characteristics is due, at leastin part, to using several slightly different (though nominallyidentical) master patterns and to a lack of automation of thereplication process.

Example 2 Sensor Fabrication

A sol-gel derived low-index nanoporous silica thin-film (See, e.g., U.S.Pat. No. 6,395,651) was incorporated into a sensor in place of theUV-cured epoxy used in previous designs. Since the low-index materialcures by heat rather than UV exposure, it was necessary to develop a newfabrication process. It was desirable to retain a low-cost imprintingmethod, though it was obvious that a plastic substrate could not sustainthe requisite high temperatures for porous glass annealing. One possibleapproach to sol-gel glass imprinting was to use a polydimethylsiloxane(PDMS) mold and a glass substrate. See, e.g., Parashar, et al.,Nano-replication of diffractive optical elements in sol-gel derivedglasses. Microelectronic Engineering, 2003. 67-8: p. 710-719.

The sub-wavelength grating structure of the low-k biosensor wasfabricated using a combination of lithography, molding, and imprintingprocesses. Sylgard 184 PDMS (Dow Corning) daughter molds are firstreplicated from a silicon master wafer patterned with a positive imageof the surface structure desired in the finished sensor. To facilitaterelease of the PDMS mold from the silicon wafer, the wafer was surfacetreated with a release layer of dimethyldichlorosilane (Repel Silane,Amersham Biosciences). See, e.g., Beck et al., Improving stamps for 10nm level wafer scale nanoimprint lithography. MicroelectronicEngineering, 2002. 61-2: p. 441-448. The PDMS replicas are then used toimprint a thin film of uncured NANOGLASS® (Honeywell Elec. Mat.), alow-index sol-gel glass, spun-on to a glass substrate. Once thelow-index dielectric becomes rigid, the flexible PDMS mold is removedand the sol-gel glass is fully cured by further baking. The sensorstructure is completed by evaporating 175 nm of TiO₂ onto the patternedsurface. A subsequent surface treatment with dimethyldichlorosilaneencourages bio-adsorption and promotes sensor stability. A schematicillustrating the cross-section of the device is shown in FIG. 7.

The polymer structure is similar to that described in a previouspublication. See, e.g., Cunningham et al., A plastic calorimetricresonant optical biosensor for multiparallel detection of label-freebiochemical interactions. Sensors and Actuators B, 2002. 85: p. 219-226.Both structures use a 550 nm period and 170 nm imprint depth, though thepolyester/polymer and low-index porous glass devices use 120 nm and 165nm TiO₂ coatings, respectively. The two devices will be referred to asthe “polymer” and “porous glass” sensors throughout the remainder of theexamples. The polymer devices were provided as an array of sensorsaligned and attached to bottomless 96-well standard microtiter plates(SRU Biosystems). The porous glass devices are fabricated on 75 mm×25mm×1 mm glass microscope slides. Adhesive rubber wells (ResearchInternational Corp.) are attached to the glass surface to provide liquidcontainment for 5-6 sensors on each slide.

Deionized water (DI H₂O, n=1.333) and Isopropyl Alcohol (IPA, n=1.378)were used to determine the bulk shift coefficient of each sensor. First,DI H₂O was pipetted onto the surface of the sensor and the PWV wasmeasured. The configuration of the readout instrument has been reportedpreviously. See, e.g., Cunningham et al., Colorimetric resonantreflection as a direct biochemical assay technique. Sensors andActuators B, 2002. 81:316-328. A broad wavelength light source wascoupled to an optical fiber that illuminates a ˜2 mm diameter region ofthe photonic crystal surface from below the substrate at normalincidence. Reflected light was collected by a second optical fiber thatis bundled next to the illuminating fiber, and measured by aspectrometer. An automated motion stage enables parallel collection ofreflectance data at timed intervals from many wells in order to acquirekinetic information.

Next, the surface was thoroughly dried and the previous step wasrepeated for IPA. The bulk shift coefficient between DI H₂O and IPA wasthen calculated as the change in PWV divided by the change in bulkrefractive index.

Example 3 PPL Bio-Adhesion Test

Sensitivity to surface-adsorbed material was characterized by thedetection of a single layer film of Poly(Lys, Phe) (PPL; Sigma-Aldrich;MW=35,400 Da) prepared to a 0.5 mg/ml solution with 0.01 M phosphatebuffered saline (PBS; pH =7.4) applied to the sensor surface. At asampling interval of 1 minute, the bio-adhesion test commenced with thepipetting of PBS into the test wells. After 10 minutes, the buffer wasreplaced with PPL solution and was allowed to stabilize for 30 minutes.The wells were then washed three times and filled with PBS for the final30 minutes of data acquisition.

PPL was deposited on 5 porous glass and 9 polymer sensors. FIG. 10compares the kinetic plots of each device, showing a ˜4× increase insurface sensitivity for the porous glass sensor. The first stepestablishes a baseline, the second corresponds to the rapid surfaceadsorption and saturation of PPL, and the final third of the curveillustrates the monolayer stability after eliminating weakly ornon-specifically bound molecules by rinsing with PBS buffer. The PWVshifts generated during PPL immobilization onto the porous glass sensorsaturate more slowly than that measured using the polymer devices.Clearly, the porous glass sensor surface is significantly less conduciveto protein monolayer adsorption. Further surface chemistry optimizationshould mitigate this effect. Nonetheless, the porous glass sensorexhibits excellent stability after unbound molecules are washed away.

Example 4 Multilayer Polymer Test

In order to characterize the differential sensitivity as a function ofdistance from the sensor surface, a series of polymer electrolytemonolayers were deposited on the sensors. By alternating betweenpositively and negatively charged polymer layers, a stack of uniform,self-limiting polymers may be formed on the sensor while it iscontinuously monitored on the detection instrument. See, e.g.,Cunningham et al., Enhancing the surface sensitivity of calorimetricresonant optical biosensors. Sensors and Actuators B, 2002.87(2):365-370. Three different polyelectrolytes were deposited onto thesensor surface: anionic poly(sodium 4-styrenesulfonate) (PSS; MW=70,000Da), cationic poly(ethyleniminie) (PEI; MW=60,000 Da), and cationicpoly(allylamine hydrochloride) (PAH; MW=70,000 Da). The polyelectrolyteswere purchased from Sigma-Aldrich. A 0.9 M NaCl buffer solution(Sigma-Aldrich) was prepared with deionized water. The polyelectrolyteswere dissolved in the buffer solution at a concentration of 5 mg/ml. Ata 1 minute sampling interval, the multilayer surface sensitivitycharacterization was performed in 5 minute steps. First, NaCl buffer waspipetted into the sensor wells. Next, the buffer was removed andreplaced by PEI solution. The wells were then washed 3 times and filledwith buffer. The previous 2 steps were repeated for PSS and PAH until 7PSS-PAH layers had been deposited atop the single PEI layer.

The 14 alternating layers of PSS and PAH described previously each causea measurable shift in the detected PWV as they are adsorbed onto thesurface. FIG. 11 gives a spatial profile of PWV shift versus polymerthickness, where each PWV shift was measured in buffer after the washstep. Each monolayer of polyelectrolyte is approximately 4.4 nm thickand has a refractive index of 1.49. See, e.g., Picart et al.,Determination of structural parameters characterizing thin films byoptical methods: A comparison between scanning angle reflectometry andoptical waveguide lightmode spectroscopy. Journal of Chemical Physics,2001. 115(2): p. 1086-1094. The porous glass sensor exhibits an averagesurface sensitivity ˜1.5× that of the polymer sensor. However, note thateach of the first 2 layers (˜9 nm) deposited onto the porous glassdevice cause a PWV shift with twice the magnitude of each of theremaining layers, while no such effect is observed for the polymerdevice.

Example 5 Bioassay: Protein A—IgG

To demonstrate selective detection by the proposed device, a bioassaywas performed that characterizes the affinity of human, sheep andchicken IgG for protein A. Protein A (Pierce Biotechnology) was preparedwith 0.01 M PBS to a concentration of 0.5 mg/ml. The buffer was filteredwith a 0.22 μm syringe filter (Nalgene) before use. Human, sheep, andchicken immunoglobulin-G (IgG) serums (Sigma-Aldrich) were diluted in0.01 M PBS to a concentration of 0.5 mg/ml. Allowing thirty minutesbetween each step and sampling at a one minute interval, PBS solutionwas first pipetted into the sensors wells. Next, the buffer was replacedby protein A solution. The well was then rinsed 3 times and filled withbuffer. After the signal stabilized, the buffer in three of the wellswas replaced by human, sheep, or chicken IgG, while the fourth was leftas a reference containing only the buffer. Finally, the IgGs wereremoved and the wells were again rinsed and filled with PBS for thefinal 30 minutes of data acquisition.

Protein A was introduced into 15 porous glass and 16 polymer sensorwells. The resulting PWV shift after the wash step was ˜4× greater forthe porous glass devices. FIG. 13 illustrates the measured bindingkinetics of human, sheep, and chicken IgG with protein A for the porousglass sensor, while FIG. 14 gives an endpoint PWV shift comparison(relative to a reference well without IgG) between the two devices foreach antibody. Protein A surface adsorption saturated much more quicklyon the polymer sensor surface, similar to that observed in the PPLbio-adhesion test. The porous glass device exhibits increasingly greatersensitivity for antibodies with higher affinity for protein A. Human IgGbinding was detected with twice the sensitivity, while Chicken IgG,lacking any specificity for protein A (See, e.g., Richman et al., Thebinding of Staphylococci protein A by the sera of different animalspecies. Journal of Immunology, 1982. 128: p. 2300-2305), results in anequivalent response and provides a measure of non-specific binding.

A photonic crystal biosensor is designed to couple electromagneticenergy to biological material deposited upon its surface from a liquidtest sample. While the device itself consists of a low refractive indexsurface structure and a high refractive index dielectric coating, theliquid test sample that fills in the surface structure must also beconsidered an integral part of the sensor—and the only dynamic componentthat can induce a change of resonant wavelength. The motivation forincorporating an extremely low refractive index material into thephotonic crystal biosensor structure is to bias the electromagneticfield of the resonant wavelength to interact more strongly with theliquid test sample and less strongly with the interior regions of thephotonic crystal that do not interact with surface adsorbed material.

The use of spin-on low-k dielectric materials leverages off the largeinvestments made in the integrated circuit manufacturing community, whorequire rapid processing, structural stability, and exclusion of liquidpenetration. A unique aspect of this work is the use of an imprintingmethod to accurately impart a submicron surface structure to ananoporous glass film without the use of photolithography. The presenceof the imprint tool on the surface of the low-k film during the initialstage of the curing process did not alter the refractive index of thefinal cured structured film. The imprinting method enables substantialcost to be incurred only in the production of the “master” siliconwafer, which is in turn used to produce a nearly unlimited number of“daughter” PDMS imprinting tools. Each PDMS tool can be used to producea large number of sensor structures without damage to the tool becauselittle force is needed to make the spun-on liquid low-k layer conform tothe tool. After imprinting, the low-k dielectric film is cured rapidlyon hotplates, using methods that are easily automated. The use of aflexible imprinting tool was found to be advantageous over imprintingfrom the silicon master wafer directly, as the PDMS mold was easier torelease from the partially cured low-k film, and was capable of allowingpermeation of volatile solvent released during the cure process.Although only 1×3-inch microscope slide regions were imprinted in thework shown here, the imprinting method can be scaled to larger surfaceareas to enable production of sensor areas large enough to cover anentire 96-well or 384-well standard microplate (approximately3×5-inches).

An interesting and useful result found during comparison of porous glasssensor structures with polymer sensor structures is the disparity insensitivity gains between bulk refractive index sensitivity andsurface-adsorbed layer sensitivity. While computer models accuratelypredict the ˜2× sensitivity increase measured for PWV shift induced by abulk refractive index change of the solution covering the porous glasssensor surface, a ˜4× increase of PWV shift was consistently measuredfor thin layers of adsorbed material. By measuring the PWV shift as afunction of thickness using the polymer multilayer experiment (FIG. 11),we are able to characterize the strength of interaction of the coupledelectromagnetic field as a function of distance away from the sensorsurface. For the porous glass sensors, the interaction is particularlystrong for the first few monolayers of adsorbed polymer, while therelationship between polymer thickness and PWV is highly linear for eachadsorbed monolayer on the polymer sensor structure. The interactionbetween the test sample and the resonant electromagnetic fielddistribution is highly complex, as detected material can adsorb to thehorizontal and vertical surfaces of the structure, where acharacteristic field profile extends into the sample from each surface.Surface-based detection sensitivity is enhanced beyond the improvementsin bulk sensitivity for the porous glass biosensor. Because the majorityof biomolecular interactions are expected to occur within the first fewnanometers from the sensor surface, the surface sensitivity is ofgreatest importance for increasing sensitivity in the context ofsurface-based biochemical assays.

1. A method of making a sensor comprising: (a) putting a liquidelastomer into a sub-wavelength period grating structure mold and curingthe liquid elastomer to a form a negative mold; (b) removing thenegative mold from the mold; (c) setting the negative mold into anuncured nanoporous film, wherein the uncured nanoporous film issupported by a substrate; (d) allowing the nanoporous film to fullycure, partially cure or not cure; (e) removing the negative mold fromthe fully cured, partially cured, or uncured nanoporous film; (f)depositing a high dielectric constant dielectric material onto the fullycured, partially cured, or uncured nanoporous film, wherein a sensor isformed.
 2. The method of claim 1, wherein the nanoporous film comprisesporous silica xerogel, porous aerogels, porous hydrogen silsesquioxane,a B staged polymer, porous methyl silsesquioxane, porous poly(aryleneether), or combinations thereof.
 3. The method of claim 1, wherein thesubstrate comprises plastic, epoxy, or glass.
 4. The method of claim 1,wherein the high dielectric constant dielectric material is tin oxide,tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, ora combination thereof.
 5. The method of claim 1, wherein the refractiveindex of the high dielectric constant dielectric material is from about1.8 to about 3.0.
 6. The method of claim 1, further comprisingdepositing a cover layer on the high dielectric constant dielectricmaterial.
 7. The method of claim 1, further comprising immobilizing oneor more specific binding substances on the surface of the highdielectric constant dielectric material.
 8. The method of claim 6,further comprising immobilizing one or more specific binding substanceson the surface of the cover layer.
 9. The method of claim 7, wherein theone or more specific binding substances do not comprise a detectionlabel.
 10. The method of claim 8, wherein the one or more specificbinding substances do not comprise a detection label.
 11. The method ofclaim 1, wherein the refractive index of the nanoporous film is fromabout 1.1 to about 2.2.
 12. The method of claim 1, wherein therefractive index of the nanoporous film is from about 1.1 to about 1.5.13. The method of claim 1, wherein the period of the sub-wavelengthperiod grating structure is from about 200 nm to about 1,500 nm and thedepth of the sub-wavelength period grating structure is from about 50 nmto about 900 nm.
 14. The method of claim 1, wherein the refractive indexof the substrate is about 1.4 to about 1.6.
 15. The method of claim 1,wherein the thickness of the high dielectric constant dielectricmaterial is about 30 nm to about 700 nm and the thickness of thenanoporous film is about 10 mn to about 5,000 nm.
 16. A method of makinga sensor comprising: (a) curing a layer of nanoporous material having alow refractive index onto a substrate; (b) depositing a high dielectricconstant dielectric material on top of the nanoporous material; and (c)patterning the high dielectric constant dielectric material into asub-wavelenth period grating structure.
 17. The method of claim 16,wherein the nanoporous material comprises porous silica xerogel, porousaerogels, porous hydrogen silsesquioxane, a B staged polymer, porousmethyl silsesquioxane, porous poly(arylene ether), or combinationsthereof.
 18. The method of claim 16, wherein the substrate comprisesplastic, epoxy, or glass.
 19. The method of claim 16, wherein the highdielectric constant dielectric material is tin oxide, tantalumpentoxide, zinc sulfide, titanium dioxide, silicon nitride, or acombination thereof.
 20. The method of claim 16, wherein the refractiveindex of the high dielectric constant dielectric material is about 1.8to about 3.0.
 21. The method of claim 16, further comprising depositinga cover layer on the high refractive index dielectric material.
 22. Themethod of claim 16, further comprising immobilizing one or more specificbinding substances on the surface of the high refractive indexdielectric material.
 23. The method of claim 21, further comprisingimmobilizing one or more specific binding substances on the surface ofthe cover layer.
 24. The method of claim 22, wherein the one or morespecific binding substances do not comprise a detection label.
 25. Themethod of claim 23, wherein the one or more specific binding substancesdo not comprise a detection label.
 26. The method of claim 16, whereinthe refractive index of the nanoporous material is from about 1.1 toabout 2.2.
 27. The method of claim 16, wherein the refractive index ofthe nanoporous material is from about 1.1 to about 1.5.
 28. The methodof claim 16, wherein the period of the sub-wavelength period gratingstructure is about 200 nm to about 1,500 nm and the depth of thesub-wavelength period grating structure is about 50 nm to about 900 nm.29. The method of claim 16, wherein the refractive index of thesubstrate is about 1.4 to about 1.6.
 30. The method of claim 16, whereinthe thickness of the high dielectric constant dielectric material isabout 30 nm to about 700 nm and the thickness of the nanoporous materialis about 10 mn to about 5,000 nm.